Single photon emission computerized tomography (SPECT) and positron emission tomography (PET) depend upon the detection of gamma ray photons emitted by, or generated as a result of, radio-pharmacological compounds introduced into the body to image regions of the body. Photon detectors used in SPECT and PET imaging systems that provide data from the emitted photons in order to produce images provide a spatial location at which a photon is detected and generally a measure of the energy of the detected photon. In PET imaging two photons are detected in coincidence and in addition to providing a spatial location for a detected photon, photon detectors also provide an indication of the time at which they detect a photon.
Traditional photon detectors for PET, SPECT and other applications, generally use an NaI(Tl) or BGO scintillation crystal to detect photons. However, NaI(Tl) has a relatively low density and low stopping power per gram of material. In order to provide detectors having good detection efficiencies for photons, NaI(Tl) crystals used in these detectors must be made relatively thick. However, as the thickness of a detector crystal increases, generally, the spatial resolution of the detector decreases. BGO scintillation crystals on the other hand exhibit low light output for detected photons and therefore generally poor energy resolution Furthermore, scintillation light in both NaI(Tl) and BGO crystals caused by the passage of a photon through the crystal lasts for a relatively long time. This complicates accurate coincidence timing using these detectors.
Semiconductor materials with high atomic numbers and relatively high densities such as CdZnTe, CdTe, HgI2, InSb, Ge, GaAs, Si, PbCs, PbS, or GaAlAs, have a high stopping power for photons per centimeter path length in the semiconductor material. Therefore, for the same stopping power, crystals made from these materials are generally thinner than detector crystals made from other materials that are used for photon detectors. These materials if they are formed into crystals of sufficient thickness can therefore be used to provide detectors with good photon detection efficiencies and improved spatial resolution. However, it difficult to manufacture thick semiconductor crystals for photon detectors using present state of the art technology.
Gamma cameras for detecting photons for SPECT, PET and other applications requiring photon detection and location, have been fabricated from the semiconductor materials listed above. Often these gamma cameras comprise arrays of pixelated detector modules, hereinafter referred to as “pixelated detectors”. A pixelated detector is described in PCT publication WO 98/23974, the disclosure of which is incorporated herein by reference.
FIGS. 1A and 1B schematically illustrate a pixelated detector and a gamma camera comprising pixelated detectors respectively. FIG. 1A shows a typical construction of a pixelated detector 20 comprising a crystal 22 formed from a semiconductor material such as one of those noted above. A face 24 of crystal 22 has a large single cathode electrode 26. An opposite face 28 of crystal 22 has an anode 30 comprising a rectangular array of identical small square anode pixels 32. Typically, sizes of anode pixels 32 vary between 1 and 4 mm2, and the thickness of crystal 22, between anode 30 and cathode 26 is on the order of millimeters to a centimeter. In operation, a voltage difference is applied between anode and cathode so that an electric field, hereinafter referred to as a “detector field”, is generated in crystal 22. This field is typically on the order of a few kilovolts per centimeter.
When a photon having an energy typical of the energies of photons used in SPECT or PET applications is incident on crystal 22, it generally loses all its energy in crystal 22 by ionization and leaves pairs of mobile electrons and holes in a small localized region of crystal 22. As a result of the detector field, the holes drift to cathode 26 and the electrons drift to anode 30, thereby inducing charges on anode pixels 32 and cathode 26. The induced charges on anode pixels 32 are sensed and generally partially processed by appropriate electronic circuits comprised in ASICs (not shown) located in a detector base 34 on which detector 20 is mounted. Detector base 34 comprises connection pins 36 for mounting to a motherboard (not shown) and transmitting signals from the ASICs to the motherboard. Signals from the induced charges on pixels 32 are used to determine the time at which a photon is detected, how much energy the detected photon deposited in the crystal and where in the crystal the photon interaction took place.
FIG. 1B shows a rectangular gamma camera 38 comprising twenty pixelated detectors 20 arranged to form a rectangular array of five rows of four detectors 20 each. Detectors 20 are shown mounted on a motherboard 39. In practice, gamma cameras comprising larger arrays of pixelated detectors are generally used.
Pixelated detectors provide reasonable spatial resolution and detection efficiency for incident photons. However, signals from anode pixels of a pixelated detector of this type are highly variable in shape and exhibit considerable jitter in signal timing with respect to the time that detected photons are incident on the detector. As a result, for applications such as PET, which require accurate coincidence measurements between different detectors, output signals from anode pixels are not readily useable. Furthermore, pixelated detectors do not provide highly accurate measurements of energies of detected photons such as are required in SPECT, PET and other applications.
The energy of a photon detected by a semiconductor detector is generally determined from an estimate of the total number of electron-hole pairs produced in the detector's crystal when the photon ionizes material of the crystal. This is generally determined from the number of electrons produced in the ionizing event, which is estimated from the charge collected on the anode of the detector. The energy resolution of the detector is a function of how accurately the number of electron-hole pairs produced in the detector by a detected photon can be estimated.
If all the electrons and holes produced by a photon detected in a semiconductor detector were collected by the detector electrodes, then the induced charge on either the anode or cathode of the detector would be an appropriate measure of the energy of the photon. However, the holes are generally subject to relatively severe charge trapping problems in the detector crystal and they drift in the crystal at approximately one tenth the velocity of the electrons. Charge trapping and the slow drift velocity of the holes degrade the accuracy with which the induced charges on either the cathode or anode of a semiconductor detector can be used to estimate the energy of a detected photon.
Semiconductor photon detectors with improved energy resolution that do not detect the location of detected photons, but only their energies, are described in “Unipolar Charge Sensing with Coplanar Electrodes—Application to Semiconductor Detectors,” by P. N. Luke, IEEE Transaction on Nuclear Science, vol. 42(4), 1995, p. 207, and in “Coplanar Grid Patterns and Their Effect on Energy Resolution of CdZnTe Detectors” by Z. He, et al, presented at the 1997 IEEE Nuclear Science Symposium and Medical Imaging Conference, Albuquerque, N. Mex., Nov. 9–15, 1997, which are incorporated herein by reference. For the described detectors, the effects of charge trapping and drift motion of holes on energy determinations of photons detected by the detectors are substantially reduced.
FIG. 2 shows a detector 40 of the type described in the referenced articles. Detector 40 comprises a semiconductor crystal 42 (CdZnTe is the semiconductor used in the referenced papers) sandwiched between a cathode 44 and an anode 46. Anode 46 comprises two coplanar comb shaped electrodes, a collecting comb anode 48 and a non-collecting comb anode 50. A voltage difference VAC is applied between cathode 44 and non-collecting anode 50 so that cathode 44 is from a few hundreds of volts to a few kilovolts negative with respect to non-collecting anode 50. A voltage VAA is applied between collecting anode 48 and non-collecting anode 50 so that collecting anode 48 is positive with respect to non-collecting anode 50 by a few tens of volts to a hundred volts.
Because of the voltage difference between collecting anode 48 and non-collecting anode 50, when a photon ionizes material in the detector crystal, collecting anode 48 collects substantially all the electrons produced by the photon. The charge on collecting anode 48 is the sum of the charge of the collected electrons and a charge induced by the electric field of trapped and slowly moving holes. The charge on non-collecting anode 50 is substantially only a charge induced by the electric field of the trapped and slowly moving holes. The charge induced by the slowly moving holes is the same for both comb electrodes. Thus, the difference between the charges on collecting anode 48 and non-collecting anode 50 is the charge on collecting anode 48 resulting from the collected electrons only. The difference between the sensed charges (assuming proper calibration and adjustment of electronic circuits used to sense and process the charges) on collecting and non-collecting anodes 48 and 50 respectively, is therefore a relatively accurate measure of the number of electrons produced by the photon and is used to determine the energy of the photon.
While a detector with coplanar comb electrodes of the type shown in FIG. 2 has improved energy resolution for detected photons it does not measure the positions of detected photons.
It would be desirable to have a pixelated detector for photons that has one or more of, improved energy resolution, stopping power and an output signal that can be used as an accurate indicator of the time at which a detected photon is incident on the detector.